The present invention is generally in the area of polymeric implants for bone repair and replacement, and particular is a polymer-hydroxyapatite composite.
Successful design of an implant to replace skeletal tissue requires knowledge of the structure and mechanical properties of bone and an understanding of the means by which grafts become incorporated into the body. This information can then be used to define desirable characteristics of the implant to ensure that the graft functions in a manner comparable to organic tissue.
The mechanical properties of bone are related to the internal organization of the material, as reviewed by Roesler, H., "The History of Some Fundamental Concepts in Bone Biomechanics," Journal of Biomechanics, 20, 1025-34 (1987). Cortical bone is classified as a material of less than 30% porosity, as described by Keaveny, T. M. and W. C. Hayes, "Mechanical Properties of Cortical and Trabecular Bone," in Bone Volume 7: Bone Growth-B, B. K. Hall, ed., Boca Raton: CRC Press, 285-344 (1992), as a "solid containing a series of voids (Haversian canals, Volkmann's canals, lacunae and canaliculi). The porosity of cortical bone tissue (typically 10%) is primarily a function of the density of these voids." In contrast, cancellous/trabecular bone is "a network of small, interconnected plates and rods of individual trabeculae with relatively large spaces between the trabeculae." Trabecular bone has a porosity of 50-90% which is a function of the space between the trabeculae.
The material properties of bone are based on determinations of the elastic modulus, compressive and tensile strengths. As a general rule, bone is stronger in compression than in tension and cortical is stronger than trabecular bone. Ranges of reported elastic modulus have been from 10 MPa to 25 GPa (10 MPa to 2 GPa for cancellous bone; 4 to 25 GPa for cortical and cancellous bone); compressive strength between 40 and 280 MPa (40 to 280 MPa for cancellous bone; 138 to 193 MPa for cortical bone); and tensile strength between 3.5 MPa and 150 MPa (3.5 to 150 MPa for cancellous bone; 69 to 133 MPa for cortical bone) (Friedlaender and Goldberg, Bone and Cartilage Allografts Park Ridge: American Academy of Orthopedic Surgeons 1991; Jarcho, "Calcium Phosphate Ceramics as Hard Tissue Prosthetics" Clin. Orthopedics and Related Research 157, 259-278 1981; Gibson, "The Mechanical Behavior of Cancellous Bone" J. Biomechan. 18(5), 317-328 1985; Keaveny and Hayes 1992).
Mechanisms by which bone may fail include brittle fracture from impact loading and fatigue from constant or cyclic stress. Stresses may act in tension, compression, or shear along one or more of the axes of the bone. A synthetic bone substitute must resist failure by any of these stresses at their physiological levels. A factor of safety on the strength of the implant may ensure that the implant will be structurally sound when subject to hyperphysiological stresses.
A graft may be necessary when bone fails and does not repair itself in the normal amount of time or when bone loss occurs through fracture or tumor. Bone grafts must serve a dual function: to provide mechanical stability and to be a source of osteogenesis. Since skeletal injuries are repaired by the regeneration of bone rather than by the formation of scar tissue, grafting is a viable means of promoting healing of osseous defects, as reviewed by Friedlaender, G. E., "Current Concepts Review: Bone Grafts," Journal of Bone and Joint Surgery, 69A(5), 786-790 (1987). Osteoinduction and osteoconduction are two mechanisms by which a graft may stimulate the growth of new bone. In the former case, inductive signals of little-understood nature lead to the phenotypic conversion of connective tissue cells to bone cells. In the latter, the implant provides a scaffold for bony ingrowth.
The bone remodeling cycle is a continuous event involving the resorption of pre-existing bone by osteoclasts and the formation of new bone by the work of osteoblasts. Normally, these two phases are synchronous and bone mass remains constant. However, the processes become uncoupled when bone defects heal and grafts are incorporated. Osteoclasts resorb the graft, a process which may take months. More porous grafts revascularize more quickly and graft resorption is more complete. After graft has been resorbed, bone formation begins. Bone mass and mechanical strength return to near normal.
Present methods for the repair of bony defects include grafts of organic and synthetic construction. Three types of organic grafts are commonly used: autografts, allografts, and xenografts. An autograft is tissue transplanted from one site to another in the patient. The benefits of using the patient's tissue are that the graft will not evoke a strong immune response and that the material is vascularized, which allows for speedy incorporation. However, using an autograft requires a second surgery, which increases the risk of infection and introduces additional weakness at the harvest site. Further, bone available for grafting may be removed from a limited number of sites, for example, the fibula, ribs and iliac crest. An allograft is tissue taken from a different organism of the same species, and a xenograft from an organism of a different species. The latter types of tissue are readily available in larger quantities than autografts, but genetic differences between the donor and recipient may lead to rejection of the graft.
Synthetic implants may obviate many of the problems associated with organic grafts. Further, synthetics can be produced in a variety of stock shapes and sizes, enabling the surgeon to select implants as his needs dictate, as described by Coombes, A. D. A. and J. D. Heckman, "Gel Casting of Resorbable Polymers: Processing and Applications," Biomaterials, 13(4), 217-224 (1992). Metals, calcium phosphate ceramics and polymers have all been used in grafting applications.
Calcium phosphate ceramics are used as implants in the repair of bone defects because these materials are non-toxic, non-immunogenic, and are composed of calcium and phosphate ions, the main constituents of bone (Jarcho, 1981; Frame, J. W., "Hydroxyapatite as a biomaterial for alveolar ridge augmentation," Int. J. Oral Maxillofacial Surgery, 16, 642-55 (1987); Parsons, et al. "Osteoconductive Composite Grouts for Orthopedic Use," Annals N.Y. Academy of Sciences, 523, 190-207 (1988)). Both tricalcium phosphate (TCP) [Ca.sub.3 (PO.sub.4).sub.2 ] and hydroxyapatite (HA) [Ca.sub.10 (PO.sub.4).sub.6 (OH.sub.2 ] have been widely studied for this reason. Calcium phosphate implants are osteoconductive, and have the apparent ability to become directly bonded to bone, as reported by Jarcho 1981. As a result, a strong bone-implant interface is created.
Calcium phosphate ceramics have a degree of bioresorbability which is governed by their chemistry and material structure. High density HA and TCP implants exhibit little resorption, while porous ones are more easily broken down by dissolution in body fluids and resorbed by phagocytosis. However, TCP degrades more quickly than HA structures of the same porosity in vitro. In fact, HA is relatively insoluble in aqueous environments.
Use of calcium phosphates in bone grafting has been investigated because of the chemical similarities between the ceramics and the mineral matrix found in the teeth and bones of vertebrates. This characteristic of the material makes it a good candidate as a source of osteogenesis. However, the mechanical properties of calcium phosphate ceramics make them ill-suited to serve as a structural element. Ceramics are brittle and have low resistance to impact loading.
Biodegradable polymers are used in medicine as suture and pins for fracture fixation. These materials are well suited to implantation as they can serve as a temporary scaffold to be replaced by host tissue, degrade by hydrolysis to non-toxic products, and be excreted, as described by Kulkarni, et al., J. Biomedical Materials Research, 5, 169-81 (1971); Hollinger, J. O. and G. C. Battistone, "Biodegradable Bone Repair Materials," Clinical Orthopedics and Related Research, 207, 290-305 (1986). Four polymers widely used in medical applications are poly(paradioxanone) (PDS), poly(lactic acid) (PLA), poly(glycolic acid) (PGA), and PLAGA copolymers. Copolymerization enables modulation of the degradation time of the material. By changing the ratios of crystalline to amorphous polymers during polymerization, properties of the resulting material can be altered to suit the needs of the application. For example, PLA is crystalline and a higher PLA content in a PLAGA copolymer results in a longer degradation time, a characteristic which may be desirable if a bone defect requires structural support for an extended period of time. Conversely, a short degradation time may be desirable if ingrowth of new tissue occurs quickly and new cells need space to proliferate within the implant.
Coombes and Heckman 1992 and Hollinger 1983 have attempted to create poly(lactide-co-glycolide) [(C.sub.3 H.sub.4 O.sub.2).sub.x (C.sub.2 H.sub.2 O.sub.2).sub.y ] implants as bone substitute. Hollinger used a PLAGA of high inherent viscosity (0.92 dl/g) prepared by a solvent-non-solvent casting method. Plugs of this material were implanted in tibial defects of Walter Reed rats, and humoral defects were created as control sites in which no polymer was implanted. Examination of the defects after sacrifice of the animals at 7, 14, 21, 28 and 42 days suggested that polymer may aid in osteoinduction in the early bone repair process. However, by 42 days, the rate of repair was equivalent in controls and experimental defect sites. Coombes and Heckman described a gel casting method for producing a three-dimensional PLAGA matrix. Success of this method, i.e., creation of a strong, rubbery gel, was dependent upon high inherent viscosity of the polymer (0.76-0.79 dl/g). Material properties of the polymer matrix through a degradation cycle were the focus of the research. The modulus of the PLAGA implant before degradation was 130 MPa, equivalent to that of cancellous bone. After eight weeks degradation in phosphate buffered saline (PBS), the strength of the material had deteriorated significantly. Moreover, the microporous structure (pores 205 .mu.m in diameter) has been shown to be too small to permit the ingrowth of cells, as reported by Friedlaender and Goldberg 1991 and Jarcho 1981. From a mechanical as well as a biological standpoint, this matrix is not ideal for use as a substitute bone graft material.
It is therefore an object of the present invention to provide a matrix, and method for making the matrix, which is useful as a bone graft.
It is a further object of the present invention to provide a matrix which provides both mechanical stability and serves as a source of osteogenesis over the time period required for bone replacement.